Systems and methods for tissue characterization using multiple aperture ultrasound

ABSTRACT

Changes in tissue stiffness have long been associated with disease. Systems and methods for determining the stiffness of tissues using ultrasonography may include a device for inducing a propagating shear wave in tissue and tracking the speed of propagation, which is directly related to tissue stiffness and density. The speed of a propagating shear wave may be detected by imaging a tissue at a high frame rate and detecting the propagating wave as a perturbance in successive image frames relative to a baseline image of the tissue in an undisturbed state. In some embodiments, sufficiently high frame rates may be achieved by using a ping-based ultrasound imaging technique in which unfocused omni-directional pings are transmitted (in an imaging plane or in a hemisphere) into a region of interest. Receiving echoes of the omnidirectional pings with multiple receive apertures allows for substantially improved lateral resolution.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is related to International Application No.PCT/US2013/027120, filed Feb. 21, 2013, which application isincorporated by reference herein.

This application is also related to the following US patentapplications: Ser. No. 11/865,501, filed Oct. 1, 2007 and titled “MethodAnd Apparatus To Produce Ultrasonic Images Using Multiple Apertures”;Ser. No. 12/760,375, filed Apr. 14, 2010, published as 2010/0262013 andtitled “Universal Multiple Aperture Medical Ultrasound Probe”; Ser. No.12/760,327, filed Apr. 14, 2010 and titled “Multiple Aperture UltrasoundArray Alignment Fixture”; Ser. No. 13/279,110, filed Oct. 21, 2011 andtitled “Calibration of Ultrasound Probes”; Ser. No. 13/272,098, filedOct. 12, 2011 and titled “Multiple Aperture Probe Internal Apparatus andCable Assemblies”; Ser. No. 13/272,105, filed Oct. 12, 2011 and titled“Concave Ultrasound Transducers and 3D Arrays”; Ser. No. 13/029,907,filed Feb. 17, 2011 and titled “Point Source Transmission AndSpeed-Of-Sound Correction Using Multi-Aperture Ultrasound Imaging”; andSer. No. 13/690,989, filed Nov. 30, 2012 and titled “Motion DetectionUsing Ping-Based and Multiple Aperture Doppler Ultrasound,” and U.S.Pat. No. 10,380,399 filed Mar. 30, 2016 and titled “Ultrasound ImagingSystems and Methods for Detecting Object Motion.”

INCORPORATION BY REFERENCE

All publications and patent applications mentioned in this specificationare herein incorporated by reference to the same extent as if eachindividual publication or patent application was specifically andindividually indicated to be incorporated by reference.

FIELD

This disclosure generally relates to imaging methods and devices fordetermining a material stiffness using a multiple aperture ultrasoundprobe to produce and track ultrasonic shear waves.

BACKGROUND

Changes in tissue stiffness have long been associated with disease.Traditionally, palpation is one of the primary methods of detecting andcharacterizing tissue pathologies. It is well known that a hard masswithin an organ is often a sign of an abnormality. Several diagnosticimaging techniques have recently been developed to provide fornon-invasive characterization of tissue stiffness.

One measure of tissue stiffness is a physical quantity called Young'smodulus, which is typically expressed in units of Pascals, or morecommonly kilo Pascals (kPa). If an external uniform compression (orstress, S) is applied to a solid tissue and this induces a deformation(or strain, e) of the tissue, Young's modulus is defined simply as theratio between applied stress and the induced strain:

E=S/e.

Hard tissues have a higher Young's modulus than soft tissues. Being ableto measure the Young's modulus of a tissue helps a physician indifferentiating between benign and malignant tumors, detecting liverfibrosis and cirrhosis, detecting prostate cancer lesions, etc.

A collection of diagnostic and imaging modalities and processingtechniques have been developed to allow clinicians to evaluate tissuestiffness using ultrasonography. These techniques are collectivelyreferred to herein as Elastography. In addition to providing informationabout tissue stiffness, some elastography techniques may also be used toreveal other stiffness properties of tissue, such as axial strain,lateral strain, Poisson's Ratio, and other common strain andstrain-related parameters. Any of these or other strain-relatedparameters may be displayed in shaded grayscale or color displays toprovide visual representations of such strain-related parameters. Suchinformation may be displayed in relation to two or three dimensionaldata.

Elastography techniques may be broadly divided into two categories,“quasi-static elastography” techniques and “dynamic elastography”techniques.

In quasi-static elastography, tissue strain is induced by mechanicalcompression of a tissue region of interest, such as by pressing againsta tissue with a probe a hand or other device. In other cases, strain maybe induced by compression caused by muscular action or the movement ofadjacent organs. Images of the tissue region of interest are thenobtained in two (or more) quasi-static states, for example, nocompression and a given positive compression. Strain may be deduced fromthese two images by computing gradients of the relative local shifts ordisplacements in the images along the compression axis. Quasi-staticelastography is analogous to a physician's palpation of tissue in whichthe physician determines stiffness by pressing the tissue and detectingthe amount the tissue yields under this pressure.

In dynamic elastography, a low-frequency vibration is applied to thetissue and the speed of resulting tissue vibrations is detected. Becausethe speed of the resulting low-frequency wave is related to thestiffness of the tissue in which it travels, the stiffness of a tissuemay be approximated from wave propagation speed.

Many existing dynamic elastography techniques use ultrasound Dopplerimaging methods to detect the speed of the propagating vibrations.However, inherent limitations in standard Doppler imaging presentsubstantial challenges when attempting to measure the desiredpropagation speed. This is at least partly because the waves of mostinterest tend to have a significant propagation component in a directionperpendicular to the direction of the initial low-frequency vibration.

As used herein, the term dynamic elastography may include a wide rangeof techniques, including Acoustic Radiation Force Impulse imaging(ARFI); Virtual Touch Tissue Imaging; Shearwave Dispersion UltrasoundVibrometry (SDUV); Harmonic Motion Imaging (HMI); Supersonic ShearImaging (SSI); Spatially Modulated Ultrasound Radiation Force (SMURF)imaging.

SUMMARY OF THE DISCLOSURE

Performing Elastography with a multiple aperture ultrasound imaging(MAUI) probe provides unique advantages over prior systems and methods.For example, in some embodiments, high resolution and high frame-rateimaging capabilities of a multiple aperture probe may be combined inorder to detect a propagating shear wave as perturbations in imageframes. In other embodiments, multiple aperture Doppler imagingtechniques may be used to determine a speed of a propagating shear wave.In some embodiments, either or both of these techniques may furtherbenefit from pixel-based imaging techniques and point-sourcetransmission techniques.

In some embodiments, an ultrasound imaging system is provided,comprising a first ultrasound transducer array configured to transmit awavefront that induces a propagating shear wave in a region of interest,a second ultrasound transducer array configured to transmit circularwaveforms into the region of interest and receive echoes of the circularwaveforms, and a signal processor configured to form a plurality ofB-mode images of the region of interest from the circular waveforms at aframe rate sufficient to detect the propagating shear wave in the regionof interest.

In some embodiments, the first ultrasound transducer array comprises anarray of phased-array elements. In other embodiments, the firstultrasound transducer array comprises an annular array of piezoelectricrings, and the signal processor is further configured to focus thewavefront at various depths by adjusting phasing delays. In anotherembodiment, the first ultrasound transducer array comprises a switchedring transducer. In yet an additional embodiment, the first ultrasoundtransducer array comprises a single piezoelectric transducer.

In some embodiments, the frame rate can be at least 500 fps, at least1,000 fps, at least 2,000 fps, or at least 4,000 fps.

In one embodiment, the signal processor is further configured tocalculate a speed of the propagating shear wave by identifying a firstposition of the shear wave in a first frame of the plurality of B-modeimages, identifying a second position of the shear wave in a secondframe of the plurality of B-mode images, determining a distance traveledby the shear wave between the first frame and the second frame,determining a time elapsed between the first frame and the second frame,and dividing the distance traveled by the time elapsed.

In some embodiments, the first frame is the result of combiningsub-images formed by echoes received by multiple elements of the secondultrasound transducer array.

In another embodiment, the signal processor is configured to identifythe propagating shear wave as a point cloud moving through the region ofinterest.

In one embodiment, the signal processor is configured to define an imagewindow identifying a section of the region of interest with acombination of zooming, panning, and depth selection.

In some embodiments, the system is configured to display acontemporaneous B-mode image of a selected image window.

A method of determining a stiffness of a tissue with ultrasound isprovided, the method comprising the steps of forming a baseline image ofa region of interest with an ultrasound imaging system, transmitting anultrasonic pulse configured to induce a propagating shear wave in theregion of interest, imaging the region of interest at a frame ratesufficient to detect the propagating shear wave to form a plurality ofimage frames of the region of interest, subtracting the baseline imagefrom at least two of the formed image frames to obtain at least twodifference frames, determining a position of the propagating shear wavein the at least two difference frames, and calculating a propagationspeed of the propagating shear wave in the region of interest from thepositions in the at least two difference frames.

In some embodiments, the method further comprises calculating a tissuestiffness of the region of interest from the propagation speed.

In one embodiment, the transmitting step comprises transmitting anultrasonic pulse with a first ultrasound transducer array, and whereinthe imaging step comprises imaging the region of interest with a secondultrasound transducer array.

In another embodiment, the forming step comprises transmitting acircular waveform from a first transmit aperture and receiving echoes ona first receive aperture.

In yet another embodiment, the imaging step comprises transmitting acircular waveform from the first transmit aperture and receiving echoesof the circular waveform with the first receive aperture.

In some embodiments, the first transmit aperture and the first receiveaperture do not include overlapping transducer elements.

In another embodiment, the frame rate is at least 500 fps, at least1,000 fps, at least 2,000 fps, or at least 4,000 fps.

In some embodiments, the method further comprises identifying thepropagating shear wave as a point cloud moving through the region ofinterest.

In another embodiment, the method further comprises displaying acontemporaneous image of the region of interest, including a lineindicating a direction of transmission of the ultrasonic pulseconfigured to induce a propagating shear wave.

In one embodiment, a method of identifying tissue edges with ultrasoundimaging is provided, comprising the steps of transmitting a firstunfocused ultrasound pulse into a tissue region of interest includingone or more tissue edges, transmitting a second unfocused ultrasoundpulse into the tissue region of interest, receiving echoes of the secondunfocused ultrasound pulse, identifying one or more speckle noisepatterns associated with the one or more tissue edges in the receivedechoes, assigning fiducial markers to the one or more tissue edges,transmitting a third unfocused ultrasound pulse into the tissue regionof interest, measuring a movement of the fiducial markers, and computinga tissue density of at least one tissue within the tissue region ofinterest.

In one embodiment, the method includes forming an image of the tissueregion of interest with the received echoes and the fiducial markers.

In some embodiments, the first unfocused ultrasound pulse has greateracoustic energy than the second unfocused ultrasound pulse. In otherembodiments, the first unfocused ultrasound pulse is transmitted withmore ultrasound transducers than the second unfocused ultrasound pulse.

In some embodiments, the speckle noise patterns are caused byreflections of the first unfocused ultrasound pulse from the one or moretissue edges.

In one embodiment, the movement comprises a distance traveled by thefiducial markers in response to the third unfocused ultrasound pulse.

In another embodiment, the movement comprises a propagation speed of thefiducial markers in response to the third unfocused ultrasound pulse.

An ultrasound imaging system is provided, comprising a first ultrasoundtransducer array configured to transmit a wavefront that induces apropagating shear wave in a region of interest, a second ultrasoundtransducer array configured to transmit circular waveforms into theregion of interest and receive echoes of the circular waveforms, and asignal processor configured to form a plurality of B-mode images of theregion of interest from the circular waveforms, the signal processorbeing further configured to identify one or more speckle patterns alonga tissue edge caused by the propagating shear wave to identify thetissue edge.

In one embodiment, the first ultrasound transducer array comprises anarray of phased-array elements.

In another embodiment, the first ultrasound transducer array comprisesan annular array of piezoelectric rings, and the signal processor isfurther configured to focus the wavefront at various depths by adjustingphasing delays.

In some embodiments, the first ultrasound transducer array comprises aswitched ring transducer. In other embodiments, the first ultrasoundtransducer array comprises a single piezoelectric transducer.

In some implementations, the frame rate is at least 500 fps, at least1,000 fps, at least 2,000 fps, or at least 4,000 fps.

In one implementation, the signal processor is configured to identifythe propagating shear wave as a point cloud moving through the region ofinterest.

In another implementation, the signal processor is configured to definean image window identifying a section of the region of interest with acombination of zooming, panning, and depth selection.

In some embodiments the system is configured to display acontemporaneous B-mode image of a selected image window.

A method of identifying tissue edges with ultrasound is provided, themethod comprising the steps of forming a baseline image of a region ofinterest with an ultrasound imaging system, transmitting an ultrasonicpulse configured to induce a propagating shear wave in the region ofinterest, transmitting a plurality of unfocused ultrasound pings intothe region of interest, forming a plurality of image frames of theregion of interest from received echoes of the plurality of unfocusedultrasound pings, subtracting the baseline image from at least two ofthe formed image frames to obtain at least two difference frames,identifying one or more speckle patterns along a tissue edge in theregion of interest caused by the propagating shear wave.

In some embodiments, the frame rate is at least 500 fps.

In one embodiment, the method further includes identifying thepropagating shear wave as a point cloud moving through the region ofinterest.

In another embodiment, the method further includes displaying acontemporaneous image of the region of interest, including a lineindicating the tissue edge.

An ultrasound imaging system is provided, comprising at least oneultrasound transducer array configured to transmit one or more unfocusedultrasound pulses into a region of interest, a second ultrasoundtransducer array configured to receive echoes of the one or moreunfocused ultrasound pulses, and a signal processor configured to form aplurality of B-mode images of the region of interest from the unfocusedultrasound pulses, the signal processor being further configured toidentify one or more speckle patterns along a tissue edge caused by theone or more unfocused ultrasound pulses to identify the tissue edge.

A method of identifying tissue edges with ultrasound imaging isprovided, comprising the steps of transmitting a first unfocusedultrasound pulse into a tissue region of interest including one or moretissue edges having a different tissue density than adjacent tissues,transmitting a second unfocused ultrasound pulse into the tissue regionof interest, receiving echoes of the second unfocused ultrasound pulse,identifying one or more fingerprint patterns in the received echoesresulting from the first unfocused ultrasound pulse and associated withthe one or more tissue edges, forming an image of the tissue region ofinterest with the received echoes, displaying the image and the one ormore tissue edges.

A method of identifying tissue edges with ultrasound imaging is alsoprovided, comprising the steps of transmitting an ultrasonic pulseconfigured to induce a propagating shear wave in a region of interest,transmitting an unfocused ultrasound pulse into the region of interest,receiving echoes of the unfocused ultrasound pulse including reflectionsfrom the propagating shear wave, computing tissue densities throughoutthe region of interest based on the reflections, selecting a tissuedifferentiation value to act as a filter between tissue density types,grouping tissue densities within the tissue differentiation value,forming an image of the region of interest with the received echoes,displaying the image, and identifying tissue edges between the groupingsof tissue densities.

In one embodiment, the systems and methods herein include a thirdultrasound transducer array that is not co-located with the firstultrasound transducer array, the third ultrasound transducer array beingconfigured to transmit a second wavefront that induces a secondpropagating shear wave into the region of interest.

BRIEF DESCRIPTION OF THE DRAWINGS

The novel features of the invention are set forth with particularity inthe claims that follow. A better understanding of the features andadvantages of the present invention will be obtained by reference to thefollowing detailed description that sets forth illustrative embodiments,in which the principles of the invention are utilized, and theaccompanying drawings of which:

FIG. 1 is a schematic illustration of one embodiment of a multipleaperture ultrasound elastography probe and a propagating shear wave in aregion of interest within a viscoelastic medium.

FIG. 2 is a schematic illustration of an embodiment of a multipleaperture ultrasound elastography probe having one shear wave initiatingtransducer array and four imaging transducer arrays.

FIG. 3 is a schematic illustration of an embodiment of a multipleaperture ultrasound elastography probe having one shear wave initiatingtransducer array and two concave curved imaging transducer arrays.

FIG. 3A is an illustration of an embodiment of a multiple apertureultrasound elastography probe having a section of a continuous concavecurved array designated as a shear-wave pulse initiating area.

FIG. 3B is an illustration of an embodiment of a multiple apertureultrasound elastography probe comprising a continuous 2D concavetransducer array configured for 3D imaging with one group of elementsdesignated as a shear-wave pulse initiating area.

FIG. 4 is a schematic illustration of an annular array which may be usedfor the shear wave initiating transducer array or one or more of theimaging transducer arrays.

FIG. 5 is a flow chart illustrating one embodiment of a high resolutionmultiple aperture imaging process.

FIG. 6 is a flow chart illustrating one embodiment of a high frame ratemultiple aperture imaging process.

FIG. 7 is a flow chart illustrating one embodiment of an elastographydata capture process.

FIG. 8 is an example of a difference frame showing perturbation causedby a propagating shear wave.

FIG. 9 is an example of a different frame showing perturbation caused bypropagating shear waves emanating from independent and offset transmitsources of the same multiple aperture probe.

FIG. 10 is an example of a different frame showing edge detection causedby reflective wavefront patterns relating to differing types of tissueand the perturbation along tissue barriers caused by a conventionalunfocused insonification from a multiple aperture probe.

FIG. 11A and FIG. 11B illustrate some cross-hatched fingerprint patternsderived from data collected using a ping-based multiple aperture imagingsystem.

FIG. 12 is a diagram outlining the steps involved with automatic tissuecharacterization, edge detection and tracking using a ping-basedmultiple aperture imaging system.

FIG. 13 is a flowchart showing the steps involved with automatic tissuecharacterization using speckle noise patterns associated with tissueedges and tracking using a ping-based multiple aperture imaging system.

DETAILED DESCRIPTION

The various embodiments will be described in detail with reference tothe accompanying drawings. References made to particular examples andimplementations are for illustrative purposes, and are not intended tolimit the scope of the invention or the claims.

In some embodiments, ultrasound imaging methods are provided in which amechanical wave having a shear component and a compression component isgenerated in a viscoelastic medium (such as biological tissue). Thespeed of the resulting shear wave propagation may be measured whileimaging the medium at a high frame-rate as the shear wave propagatesthrough the medium. Speed of the propagating shear may be determined byidentifying the changing position of the shear wave in a plurality offrames obtained at known time intervals. As will be described in furtherdetail below, various embodiments of ping-based and multiple apertureultrasound imaging are particularly well-suited to obtaining highresolution and high frame-rate images for performing accurate analysisof tissue stiffness using these methods. In some embodiments aqualitative and/or quantitative analysis of received echo data may beperformed to identify regions of different hardness as compared with therest of the viscoelastic medium.

Embodiments herein provide systems and methods for performing ultrasoundelastography to determine the shear modulus of a tissue. In someembodiments, a method of determining a shear modulus comprisestransmitting a mechanical shear wave into a test medium, then imagingthe test medium using a high frame rate B-mode ultrasound imagingtechnique as the shear wave propagates through the medium. By comparingeach image frame taken during propagation of the shear wave with areference image generated prior to transmitting the shear wave, apropagation velocity may be determined.

In addition to characterizing tissue stiffness, ping based multipleaperture imaging provides the advantages for detecting tissue edges,tissues characterization, tissue movement, and, or substances within anarea of interest in the medium. Embodiments of ping based multipleaperture ultrasound systems described herein may provide variousadvantages that cannot be met by other available systems. Suchadvantages may include determining the edges of tissue structures andthe edges of tissue disparities there-in. Tissue edges can be traced atframe rates of up to 10,000 frames/second, motion detection latencies ofunder 10 ms, and the ability to detect and track tissue edges withprecision on a scale far less than a wavelength of ultrasound used.Techniques described herein may be used to detect and track motion ofpoints smaller than any resolvable object in the ultrasound imagingsystem being used.

For example, systems and methods herein may be used to detect and trackmovement of an object edge measuring greater than 0.05 mm, with lessthan 1 millisecond of reporting latency, at update rates of more than 10kHz. Position and velocity of moving objects may be tracked in sixdegrees of freedom (e.g., linear movement in X, Y, Z directions androtation about pitch, roll, and yaw axes). In some cases, systems andmethods described herein can perform even better than these measures.

The Rayleigh criterion is the generally accepted criterion fordetermining the size of a minimum resolvable detail (in terms of lateralresolution) achievable by an imaging system. The imaging process is saidto be “diffraction-limited” when the first diffraction minimum of theimage of one source point coincides with the maximum of another. TheRayleigh criterion, simplified for the case of an ultrasound imagingprobe, indicates that the size (‘r’) of the minimum resolvable detail inlateral resolution of an ultrasound probe with a total aperture of D isr≈1.22λ/D (where λ is the speed-of-ultrasound in the imaged mediumdivided by ultrasound frequency).

Because there is no transmit beamforming in a ping-based ultrasoundimaging system, there is also no axial resolution in the traditionalsense attributed to conventional phased array ultrasound. However, theterm ‘axial resolution’ is used in the traditional sense here because itconveys a somewhat similar concept: the ability to distinguish tworeflectors lying close together along a radial line originating at apoint-source transmitter. The axial resolution of a ping-basedultrasound imaging system is approximately equal to the wavelength (λ)of ultrasound being used (i.e., the speed-of-ultrasound in the imagedmedium divided by ultrasound frequency) multiplied by the number ofcycles transmitted in each ping. It is possible using ping-based imagingthat lateral and axial resolution may be significantly below themeasurements normally associated with conventional monitors on imagingsystems, and in some embodiments may not be visually distinguishable toan operator. Even though operating in this non-visual range, ping-basedmultiple aperture imaging systems can still identify, define and tracktissue data. These non-visual applications will be discussed herein.

The various motion detection and motion tracking systems and methodsdescribed herein generally utilize an imaging technique referred toherein as “ping-based multiple aperture imaging” (PMA) or “ping-basedimaging.” This disclosure is organized with a description of ping-basedimaging techniques, followed by a description of edge detectiontechniques, which in turn is followed by a description of varioushardware elements that may be used in combination with the processes andtechniques described herein.

Although the various embodiments are described herein with reference toimaging and evaluating stiffness of various anatomic structures, it willbe understood that many of the methods and devices shown and describedherein may also be used in other applications, such as imaging andevaluating non-anatomic structures and objects. For example, theultrasound probes, systems and methods described herein may be adaptedfor use in non-destructive testing or evaluation of various mechanicalobjects, structural objects or materials, such as welds, pipes, beams,plates, pressure vessels, layered structures, soil, earth, concrete,etc. Therefore, references herein to medical or anatomic imagingtargets, tissues, or organs are provided merely as non-limiting examplesof the nearly infinite variety of targets that may be imaged orevaluated using the various apparatus and techniques described herein.

Introduction to Key Terms & Concepts

As used herein the terms “ultrasound transducer” and “transducer” maycarry their ordinary meanings as understood by those skilled in the artof ultrasound imaging technologies, and may refer without limitation toany single component capable of converting an electrical signal into anultrasonic signal and/or vice versa. For example, in some embodiments,an ultrasound transducer may comprise a piezoelectric device. In otherembodiments, ultrasound transducers may comprise capacitivemicromachined ultrasound transducers (CMUT).

Transducers are often configured in arrays of multiple individualtransducer elements. As used herein, the terms “transducer array” or“array” generally refers to a collection of transducer elements mountedto a common backing plate. Such arrays may have one dimension (1D), twodimensions (2D), 1.X dimensions (e.g., 1.5D, 1.75D, etc.) or threedimensions (3D) (such arrays may be used for imaging in 2D, 3D or 4Dimaging modes). Other dimensioned arrays as understood by those skilledin the art may also be used. Annular arrays, such as concentric circulararrays and elliptical arrays may also be used. An element of atransducer array may be the smallest discretely functional component ofan array. For example, in the case of an array of piezoelectrictransducer elements, each element may be a single piezoelectric crystalor a single machined section of a piezoelectric crystal.

As used herein, the terms “transmit element” and “receive element” maycarry their ordinary meanings as understood by those skilled in the artof ultrasound imaging technologies. The term “transmit element” mayrefer without limitation to an ultrasound transducer element which atleast momentarily performs a transmit function in which an electricalsignal is converted into an ultrasound signal. Similarly, the term“receive element” may refer without limitation to an ultrasoundtransducer element which at least momentarily performs a receivefunction in which an ultrasound signal impinging on the element isconverted into an electrical signal. Transmission of ultrasound into amedium may also be referred to herein as “insonifying.” An object orstructure which reflects ultrasound waves may be referred to as a“reflector” or a “scatterer.”

As used herein, the term “aperture” may refer to a conceptual “opening”through which ultrasound signals may be sent and/or received. In actualpractice, an aperture is simply a single transducer element or a groupof transducer elements that are collectively managed as a common groupby imaging control electronics. For example, in some embodiments anaperture may be a physical grouping of elements which may be physicallyseparated from elements of an adjacent aperture. However, adjacentapertures need not necessarily be physically separated.

It should be noted that the terms “receive aperture,” “insonifyingaperture,” and/or “transmit aperture” are used herein to mean anindividual element, a group of elements within an array, or even entirearrays with in a common housing, that perform the desired transmit orreceive function from a desired physical viewpoint or aperture. In someembodiments, such transmit and receive apertures may be created asphysically separate components with dedicated functionality. In otherembodiments, any number of send and/or receive apertures may bedynamically defined electronically as needed. In other embodiments, amultiple aperture ultrasound imaging system may use a combination ofdedicated-function and dynamic-function apertures.

As used herein, the term “total aperture” refers to the total cumulativesize of all imaging apertures. In other words, the term “total aperture”may refer to one or more dimensions defined by a maximum distancebetween the furthest-most transducer elements of any combination of sendand/or receive elements used for a particular imaging cycle. Thus, thetotal aperture is made up of any number of sub-apertures designated assend or receive apertures for a particular cycle. In the case of asingle-aperture imaging arrangement, the total aperture, sub-aperture,transmit aperture, and receive aperture will all have the samedimensions. In the case of a multiple aperture imaging arrangement, thedimensions of the total aperture includes the sum of the dimensions ofall send and receive apertures.

In some embodiments, two apertures may be located adjacent one anotheron a continuous array. In still other embodiments, two apertures mayoverlap one another on a continuous array, such that at least oneelement functions as part of two separate apertures. The location,function, number of elements and physical size of an aperture may bedefined dynamically in any manner needed for a particular application.Constraints on these parameters for a particular application will bediscussed below and/or will be clear to the skilled artisan.

Elements and arrays described herein may also be multi-function. Thatis, the designation of transducer elements or arrays as transmitters inone instance does not preclude their immediate redesignation asreceivers in the next instance. Moreover, embodiments of the controlsystem herein include the capabilities for making such designationselectronically based on user inputs, pre-set scan or resolutioncriteria, or other automatically determined criteria.

Inducing Shear Waves

The propagation velocity of shear waves in tissue is related to thestiffness (Young's modulus or shear modulus) and density of tissue bythe following equation:

E=3ρ·c ²

where c is the propagation velocity of shear wave, E is Young's modulus,and ρ is the tissue density. Because the density of tissues tends tovary minimally, and because the speed term is squared, elasticity may becalculated by assuming an approximate density value and measuring onlythe speed of shear wave propagation. In some cases, the assumed densityvalue may vary depending on known information about the tissue beingimaged, such as an approximate range of densities for known organtissues. For example, liver tissue may have a density of approximately1.05 kg/l, heart tissue may be about 1.03 kg/l, and skeletal muscletissue may be about 1.04 kg/l. Variations in tissue elasticity are knownto be associated with various disease states. Therefore, cancers orother pathological conditions may be detected in tissue by measuring thepropagation velocity of shear waves passing through the tissue.

In some embodiments, a shear wave may be created within tissue byapplying a strong ultrasound pulse to the tissue. In some embodiments,the shear wave generating ultrasound pulse (also referred to herein asan “initiating” pulse or an “init” pulse) may exhibit a high amplitudeand a long duration (e.g., on the order of 100 microseconds). Theultrasound pulse may generate an acoustic radiation force to push thetissue, thereby causing layers of tissue to slide along the direction ofthe ultrasound pulse. These sliding (shear) movements of tissue may beconsidered shear waves, which are of low frequencies (e.g., from 10 to500 Hz) and may propagate in a direction perpendicular to the directionof the ultrasound pulse.

The propagation speed of a shear wave is typically on the order of about1 to 10 m/s (corresponding to tissue elasticity from 1 to 300 kPa).Consequently, a propagating shear wave may cross a 6 cm wide ultrasoundimage plane in about 6 to 60 milliseconds. Thus, in order to collect atleast three images of a fast-moving shear waves in a 6 cm wide image, aframe rate of at least 500 frames per second may be required. Mostcurrent radiology ultrasound systems refresh a complete image only every17 to 33 milliseconds (corresponding to frame rates of about 58 to about30 frames per second), which is too slow to image a propagating shearwave because the shear wave will have disappeared from the field of viewbefore a single frame can be acquired. In order to capture shear wavesin sufficient detail, frame rates of a thousand or more images persecond are needed.

Ping Based Multiple Aperture Ultrasound Imaging systems provide theability to track and measure shear waves along multiple axes around thetissue and region of interest. An organ or section of tissue that istortuous (e.g. pancreas, bowel, pulmonary vessels), can be imaged andelastographic measurement made in total because Ping Based MultipleAperture Ultrasound Imaging systems do not require the tissue to be in asingle plan. Ping transmissions can be from independent coherentsections of either a 2D probe or a 3D probe and be located in plane orout of plane. Multiple Aperture receivers collecting data out of planecontinue to have the ability to track fiducials, beamform images andmeasure phase shifts, in several embodiments the transmitter is notlocated in plane. In other embodiments, the transmitter can bephysically located in plane but may be physically located closer to, orfurther away from, the target tissue. In conventional phased arrayultrasound, such asymmetries in probe shape could easily cause phasedisturbance in collecting imaging data. PMA systems, however, do notrequire transmissions from a symmetric array or probe in order toaccommodate phase shift and subsequently can provide off axis imagingand tissue density information.

High Frame Rate Ultrasound Imaging

The frame rate of a scanline-based ultrasound imaging system is thepulse-repetition frequency (PRF, which is limited by the round-triptravel time of ultrasound in the imaged medium) divided by the number ofscanlines per frame. Typical scanline-based ultrasound imaging systemsuse between about 64 and about 192 scanlines per frame, resulting intypical frame rates of only about 50 frames per second.

By using ping-based ultrasound imaging techniques, some ultrasoundimaging systems and methods are capable of achieving frame rates on theorder of thousands of frames per second. Some embodiments of suchsystems and methods are able to obtain an entire 2D image from a singletransmit pulse, and can achieve a pulse rate (and therefore, a framerate) of 4000 per second or higher when imaging to a depth of 18 cm.With this refresh rate it is possible to capture a shear wave atincrements of about 2.5 mm of travel for the fastest waves, and evenshorter increments for slower shear waves. When imaging at shallowerdepths, even higher frame rates may be achieved. For example, whenimaging at a depth of 2 cm, a ping-based ultrasound imaging system mayachieve a pulse rate (and therefore, a frame rate) of about 75,000frames per second. Still higher frame rates may be achieved bytransmitting overlapping pulses or pings (e.g, as described below).

In contrast to conventional scanline-based phased array ultrasoundimaging systems, some embodiments of multiple aperture ultrasoundimaging systems may use point source transmission during the transmitpulse. An ultrasound wavefront transmitted from a point source (alsoreferred to herein as a “ping” or an unfocused ultrasound wavefront)illuminates the entire region of interest with each circular orspherical wavefront. Echoes received from a single ping received by asingle receive transducer element may be beamformed to form a completeimage of the insonified region of interest. Combining data and imagesfrom multiple receive transducers across a wide probe, and combiningdata from multiple pings, very high resolution images may be obtained.Moreover, such a system allows for imaging at a very high frame rate,since the frame rate is limited only by the ping repetitionfrequency—i.e., the inverse of the round-trip travel time of atransmitted wavefront travelling between a transmit transducer element,a maximum-depth reflector, and a furthest receive transducer element. Insome embodiments, the frame rate of a ping-based imaging system may beequal to the ping repetition frequency alone. In other embodiments, ifit is desired to form a frame from more than one ping, the frame rate ofa ping-based imaging system may be equal to the ping repetitionfrequency divided by the number of pings per frame.

As used herein the terms “point source transmission” and “ping” mayrefer to an introduction of transmitted ultrasound energy into a mediumfrom a single spatial location. This may be accomplished using a singleultrasound transducer element or combination of adjacent transducerelements transmitting together. A single transmission from saidelement(s) may approximate a uniform spherical wave front, or in thecase of imaging a 2D slice it creates a uniform circular wave frontwithin the 2D slice. In some cases, a single transmission of a circularor spherical wave front from a point source transmit aperture may bereferred to herein as a “ping” or a “point source pulse” or an“unfocused pulse.” Using a number of transducer elements together canprovide unfocused wavefront. In some embodiments, unfocused wavefrontsare created using less than five transducer elements together. In someembodiments, unfocused wavefronts are created using less than ninetransducer elements together. In some embodiments, unfocused wavefrontsare created using less than thirteen transducer elements together.

Point source transmission are not restricted to transmitting throughsoft tissue or from a single direction. In some embodiments, pointssource transmissions are made directly through bone or gas filled areaslike lung. In other embodiments, point source transmissions are madethrough soft tissues and shear wave patterns are tracked through othertypes of tissue.

Point source transmissions can be made from any location or axis on aconcave, symmetric or asymmetric transducers. In some embodiments, pointsource transmissions are made in plane with the receiver transducerselements. In some embodiments, the point sources transmission are madeout of plane with the receiver transducer elements.

Point source transmission differs in its spatial characteristics from ascanline-based “phased array transmission” or a “directed pulsetransmission” which focuses energy in a particular direction (along ascanline) from the transducer element array. Phased array transmissionmanipulates the phase of a group of transducer elements in sequence soas to strengthen or steer an insonifying wave to a specific region ofinterest.

In some embodiments, multiple aperture imaging using a series oftransmit pings may operate by transmitting a point-source ping from afirst transmit aperture and receiving echoes of the transmitted pingwith elements of two or more receive apertures. A complete image may beformed by triangulating the position of reflectors based on delay timesbetween transmission and receiving echoes. As a result, each receiveaperture may form a complete image from echoes of each transmitted ping.In some embodiments, a single time domain frame may be formed bycombining images formed from echoes received at two or more receiveapertures from a single transmitted ping. In other embodiments, a singletime domain frame may be formed by combining images formed from echoesreceived at one or more receive apertures from two or more transmittedpings. In some such embodiments, the multiple transmitted pings mayoriginate from different transmit apertures.

“Beamforming” is generally understood to be a process by which imagingsignals received at multiple discrete receptors are combined to form acomplete coherent image. The process of ping-based beamforming isconsistent with this understanding. Embodiments of ping-basedbeamforming generally involve determining the position of reflectorscorresponding to portions of received echo data based on the path alongwhich an ultrasound signal may have traveled, an assumed-constant speedof sound and the elapsed time between a transmit ping and the time atwhich an echo is received. In other words, ping-based imaging involves acalculation of distance based on an assumed speed and a measured time.Once such a distance has been calculated, it is possible to triangulatethe possible positions of any given reflector. This distance calculationis made possible with accurate information about the relative positionsof transmit and receive transducer elements and the speed-of-ultrasoundin the imaged medium. As discussed in Applicants' previous applicationsreferenced above, multiple aperture and other probes may be calibratedto determine the acoustic position of each transducer element to atleast a desired degree of accuracy, and such element positioninformation may be digitally stored in a location accessible to theimaging or beamforming system.

FIG. 1 schematically illustrates one embodiment of a multiple apertureultrasound probe 10 configured for performing elastography. The probe 10of FIG. 1 includes two imaging transducer arrays 14, 16 and one shearwave initiating transducer array, which is referred to herein as an“init” transmit transducer array 12. An init transducer array may beconfigured for transmitting a relatively low frequency shear-waveinitiating pulse (also referred to herein as an “init pulse”).

The probe 10 may also be configured to be connected to an electroniccontroller 100 configured to electronically control transmitted andreceived ultrasonic signals. The controller may be configured totransmit phased array or ping ultrasound signals, to receive and processechoes received by the imaging transducer arrays, to perform a receivebeamforming process, and to form B-mode images from the received andprocessed echoes. The controller 100 may also be configured to controltransmission of shear wavefronts from the init array, and may beconfigured determine a position of a shear wave and an elasticity oftissue in a region of interest according to any of the embodimentsdescribed herein. The controller 100 may also be configured to controlimage formation, image processing, echo data storage, or any otherprocess, including the various methods and processes described herein.In some embodiments, some or all of the controller 100 can beincorporated into the probe. In other embodiments, the controller can beelectronically coupled to the probe (e.g., by a wired or wirelesselectronic communication method), but physically separate from the probeitself. In still further embodiments, one or more separate additionalcontrollers may be electronically connected to the probe 10 and/or tothe controller 100. Such additional controllers may be configured toexecute any one or more of the methods or processes described herein.

In the embodiment illustrated in FIG. 1 , the init transducer array 12is located centrally in between left 14 and right 16 lateral imagingtransducer arrays. In alternative embodiments, an init array may belocated in any other position, such as the left position 14, the rightposition 16 or another position in addition to those shown in FIG. 1 .In further embodiments, any one of several transducer arrays in amultiple aperture probe may be temporarily or permanently assigned andcontrolled to operate as an init array.

In further embodiments, an init transducer need not necessarily be aseparate array. Rather, a single transducer element or a group oftransducer elements that are part of a larger array that may otherwisebe used for imaging may be temporarily or permanently designated andcontrolled/operated as an init array.

As will be discussed in further detail below, the imaging transducerarrays 14, 16 of the probe 10 may be used for imaging the region ofinterest 50. The imaging transducer arrays 14, 16 may comprise anytransducer array construction suitable for ultrasound imaging, such as1D, 1.XD, 2D arrays of piezoelectric crystals, Capacitive MicromachinedUltrasound Transducer (CMUT) elements or Piezoelectric MicromachinedUltrasound Transducer (PMUT) elements.

Embodiments of multiple aperture ultrasound imaging probes may includeany number of imaging apertures in a wide range of physicalarrangements. For example, FIG. 2 illustrates an embodiment of amultiple aperture elastography probe 11 comprising a central inittransducer array 12 and two pairs of imaging arrays 14, 15, 16, 17 allfour of which may be used in a multiple aperture imaging process. Insome embodiments, the init array 12 may alternatively be in the positionof any of the other arrays 14, 15, 16, 17.

In some embodiments, multiple aperture probes may have a generallyconcave tissue-engaging surface, and may include a plurality of imagingapertures. In some embodiments, each individual aperture of a multipleaperture probe may comprise a separate and distinct transducer array. Inother embodiments, individual apertures may be dynamically and/orelectronically assigned on a large continuous transducer array.

FIG. 3 illustrates an embodiment of a multiple aperture elastographyprobe comprising a central init transducer array 12 and a pair ofconcave curved lateral imaging arrays 18, 20. In some embodiments,multiple imaging apertures may be dynamically assigned on one or both ofthe concave lateral arrays 18, 20 as described in Applicants' prior U.S.patent application Ser. No. 13/272,105, which is incorporated herein byreference. Alternatively, each of the concave curved lateral arrays maybe treated as a separate aperture.

FIG. 3A illustrates an embodiment of a multiple aperture elastographyprobe comprising a single continuous concave curved transducer array 19.As with other embodiments discussed above, any portion of the continuouscurved array 19 may be temporarily or permanently configured,designated, and controlled/operated as an init array.

FIG. 3B illustrates an embodiment of a multiple aperture elastographyprobe comprising a 3D array 25 as described in Applicants' priorapplication Ser. No. 13/272,105. A group of transducer elements 12 isshown designated as a shear wave initiating region. As with the aboveembodiments, any other region of the 3D array 25 may be designated as aninit region.

In some embodiments, a probe with at least three arrays may be adaptedfor elastography by replacing at least one transducer array with a lowfrequency init transducer array. In some embodiments, an init transducerarray of a multiple aperture probe may be positioned between at leasttwo other arrays. Such probe configurations may include adjustableprobes, cardiac probes, universal probes, intravenous ultrasound (IVUS)probes, endo-vaginal probes, endo-rectal probes, transesophageal probesor other probes configured for a particular application.

Similarly, any other multiple aperture or single-aperture ultrasoundimaging probe may be adapted for use with the elastography systems andmethods described herein. In still further embodiments, an init arraymay be provided on a separate probe entirely independent of an imagingprobe. For example, an init probe may be provided with a separatehousing from the housing of the imaging probe. In some embodiments, anindependent init probe may be configured to be temporarily attached toan imaging probe. In such embodiments, such a separate init probe may becontrolled by the same ultrasound imaging system as an imaging probe, orthe init probe may be controlled independently of the imaging system. Anindependently-controlled elastography init pulse controller may besynchronized with an ultrasound imaging system in order to provide theimaging system with accurate timing information indicating the time atwhich an init pulse is transmitted.

In alternative embodiments, similar frame rates may be achieved bytransmitting a plane wave front (e.g., by transmitting simultaneouspulses from several transducers in a common array), receiving echoes,and mapping the received echoes to pixel locations using techniquessimilar to those described above. Some embodiments of such plane-wavetransmitting systems may achieve frame rates similar to those achievedwith ping-based imaging techniques.

Embodiments of Shear-Wave Initiating Transducers

Regardless of probe construction, embodiments of an init array 12 may beconfigured to transmit shear-wave initiating ultrasound pulses withfrequencies between about 1 MHz and about 10 MHz. In other embodiments,the init array 12 may be configured to transmit shear-wave initiatingultrasound pulses with a frequency up to about 18 MHz or higher. In someembodiments, an ultrasound frequency for producing init pulses may beabout half of an ultrasound frequency used for imaging. Depending onmaterials and construction, a single transducer array may be capable ofproducing both low frequency ultrasound pulses for an init pulse andrelatively high frequency ultrasound pulses for imaging. However, insome embodiments it may be desirable to use transducers optimized for arelatively narrow frequency range to allow for more efficient control ofan init pulse or an imaging pulse.

Thus, in some embodiments, an init transducer array 12 may comprise aseparate array configured to function exclusively as an init array, suchas by being optimized to function efficiently within an expected initfrequency range. As a result, in some embodiments an init array may bestructurally different than separate imaging arrays. In otherembodiments an init array may be physically identical to an imagingarray, and may differ only in terms of its operation and use.

In some embodiments, the init transducer array 12 may comprise arectangular or otherwise shaped array (e.g., a 1D, 1.xD, 2D or otherrectangular array) of piezoelectric elements. In other embodiments, theinit transducer array 12 may comprise a rectangular or otherwise shapedarray of capacitive micro-machined ultrasound transducer (CMUT)elements.

In other embodiments, the init array 12 may comprise an annular array 30as shown for example in FIG. 4 . An annular array may comprise aplurality of transducer elements arranged in concentric circular orelliptical patterns. Such annular arrays 20 may also use any suitabletransducer material. In some embodiments, an init array 12 may comprisea switched ring annular transducer array.

In some embodiments, a switched-ring annular array may include adish-shaped ultrasonic transducer (e.g., a segment of a sphere) whichmay be divided into a plurality of concentric annular transducerelements of which the innermost element may be either a planar annulusor a complete dish. In some embodiments, the curvature of the frontsurface of the annular array 20 and any lens or impedance matching layerbetween the transducer and the region of interest surface may at leastpartially determine the focal length of the transducer. In otherembodiments, an annular array may be substantially planar and anacoustic lens may be employed to focus the transmitted ultrasoundenergy.

An annular array 20 may include any number of rings, such as three ringsin addition to the center disc as shown in FIG. 4 . In otherembodiments, an annular array may include 2, 4, 5, 6. 7. 8, 9, 10 ormore rings in addition to a center disc or dish. In some embodiments,the rings may be further decoupled by etching, scribing, completecutting or otherwise dividing the rings into a plurality of ringelements within each ring. In some embodiments, an annular arraytransducer for operating to depths of 25 cm may have a diameter of 40 mmwith the outer ring may have a width of approximately 1.85 mm, providinga surface area of 222 mm²; the inner ring may have a width ofapproximately 0.8 mm and lying at an approximate radius of 10.6 mm toprovide a surface area of 55 mm².

In some embodiments, each ring (or each ring element within a ring) mayhave individual electrical connections such that each ring (or ringelement) may be individually controlled as a separate transducer elementby the control system such that the rings may be phased so as to directa shear-wave initiating pulse to a desired depth within the region ofinterest. The amplitude of the energy applied may determine theamplitude of the emitted ultrasonic waves which travel away from theface of the annular array 20.

In some embodiments the size and/or number of elements in an init arraymay be determined by the shape or other properties of the shear waves tobe produced.

In some embodiments, a shear-wave initiating pulse produced by an inittransducer array 12 may be focused during transmission to providemaximum power at the region of interest. In some embodiments, the initpulse may be focused on an init line 22 (e.g., as shown in FIGS. 1, 2and 3 ). The init pulse may further be focused at a desired depth toproduce a maximum disruptive power at the desired depth. In someembodiments, the axial focus line and the focused depth point may bedetermined by transmitting pulses from a plurality of transducerelements at a set of suitable delays (i.e., using “phased array”techniques). In some embodiments, transmit delays may be omitted whenusing an annular array with a series of switched rings as discussedabove.

In some embodiments, the init pulse need not be electronicallysteerable. In such embodiments, the probe may be configured to alwaystransmit an init pulse along a consistent line relative to the probe. Insome embodiments, the expected line of the init pulse may be displayedon the ultrasound display (e.g., overlaying a contemporaneous B-modeimage of the region of interest) so as to provide an operator with avisual indication of the path of the init pulse relative to the imagedregion of interest. In such embodiments, a sonographer may manipulatethe probe until the display shows a representative init line passingthrough an object to be evaluated by elastography.

In alternative embodiments, an init pulse may be electronically steeredin a direction indicated by an operator. In such embodiments, the lineof the init pulse may be selected by an operator through any appropriateuser interface interaction without the need to move the probe. In someembodiments, the user interface interaction may include a visual displayof the init line on a display screen (e.g., overlaying a contemporaneousB-mode image of the region of interest). Once a desired init pulsedirection is chosen, an init pulse may be electronically steered so asto travel along the selected line.

Embodiments for Detecting Shear Wave Propagation Rate

Returning to FIG. 1 , an example of shear wave propagation will bedescribed. A shear wave may be initiated in a region of interest 50 froman init pulse from a multiple aperture elastography probe 10 (or anyother suitably configured elastography probe). As discussed above, theinit pulse may be focused along a line 22 extending from the inittransducer array 12 into the region of interest to at least a desireddepth. In some embodiments, the line 22 may be perpendicular to the inittransducer array 12. An initial pulse 52 transmitted along the init line22 will tend to induce a wave front 56 propagating outwards from theline 22 within the image plane. The propagating wavefront 56 induced bythe init pulse will push the tissue in the direction of propagation. Anelastic medium such as human tissue will react to this push by arestoring force which induces mechanical waves including shear waveswhich propagate transversely from the line 22 in the tissue.

Embodiments of elastographic imaging processes will now be describedwith reference to the probe construction of FIG. 1 and the flow chartsof FIGS. 5-7 . These processes may be used with any suitably configuredprobe as described above. In some embodiments, the left and rightlateral transducer arrays 14, 16 may be used to image the region ofinterest 50 with either, both or a combination of a high frame rateultrasound imaging technique and a high resolution multiple apertureultrasound imaging technique. These techniques are summarized below, andfurther details of these techniques are provided in U.S. patentapplication Ser. No. 13/029,907, which illustrates embodiments ofimaging techniques using transmission of a circular wavefront and usingreceive-only beamforming to produce an entire image from each pulse or“ping” (also referred to as ping-based imaging techniques).

The terms “high resolution imaging” and “high frame rate imaging” areused herein as abbreviated names for alternative imaging processes.These terms are not intended to be limiting or exclusive, as the “highresolution imaging” process may also be operated at a high frame raterelative to other imaging techniques, and the “high frame rate imaging”process may also produce images of a substantially higher resolutionthan other imaging techniques. Furthermore, the rate of shear wavepropagation may be detected using high frame rate imaging techniquesand/or high resolution imaging techniques other than those described orreferenced herein.

FIG. 5 illustrates an embodiment of a high resolution multiple apertureimaging process 60 that may use a multiple aperture ultrasound imagingprobe such as that shown in FIG. 1 . In some embodiments, one or both ofthe imaging arrays 14, 16 may include one or more transducer elementstemporarily or permanently designated as transmit elements T1 throughTn. The remaining transducer elements of one or both of the imagingarrays 14, 16 may be designated as receive elements.

In some embodiments, a high resolution multiple aperture ultrasoundimaging process 60 may comprise transmitting a series of successivepulses from a series of different transmit apertures (T1 . . . Tn) 62,receiving echoes 64 from each pulse with a plurality of elements on areceive aperture, and obtaining a complete image 66 from echoes receivedfrom each transmit pulse. These images may then be combined 68 into afinal high-resolution image. Embodiments of such a high resolutionmultiple aperture imaging process may be substantially similar to theprocess shown and described in Applicants' prior U.S. patent applicationSer. No. 13/029,907 referenced above.

As indicated in FIG. 5 , during a first cycle of a high resolutionimaging process, the steps of transmitting an ultrasound signal 62A,receiving echoes 64A, and forming an image 66A may be performed using afirst transmit transducer T1. During a second cycle, signals may betransmitted 62B from a different transmit transducer Ti, echoes may bereceived 64B, and a second image may be formed 66B. The process of steps62 x-66 x may be repeated using n different transmit transducers whichmay respectively be located at any desired position within an ultrasoundprobe. Once a desired number of image (also referred to as image layers)have been formed, such image layers may be combined 68 into a singleimage frame, thereby improving image quality. If desired, the process 60may then be repeated to obtain multiple time-domain frames which maythen be consecutively displayed to a user.

FIG. 6 illustrates an embodiment of a high frame rate imaging process70. In some embodiments, a high frame rate ultrasound imaging process 70may comprise transmitting successive pings from a single transmitaperture Tx 72, forming a complete image 76 from echoes received 74 fromeach transmitted ping 72, and treating each image 76 as a successivetime domain frame. In this way, slight changes in the position ofreflectors in the region of interest 50 can be sampled at a very highframe rate.

As indicated in FIG. 6 , during a first cycle, a ping may be transmittedfrom a chosen transmit transducer Tx 72A, echoes may be received 74A anda first frame may be formed 76A. The same cycle of steps transmitting72B and receiving 74B may then be repeated to produce a second frame76B, a third frame (steps 72C, 74C, 76C), and as many subsequent framesas desired or needed as described elsewhere herein.

In some embodiments, a maximum frame rate of an imaging system usingping-based imaging techniques may be reached when a ping repetitionfrequency (i.e., the frequency at which successive pings aretransmitted) is equal to an inverse of the round trip travel time (i.e.,the time for an ultrasound wave to travel from a transmit transducer toa reflector at a desired distance from the transducer, plus the time foran echo to return from the reflector to a receive transducer along thesame or a different path). In other embodiments, overlapping pings maybe used with coded excitation or other methods of distinguishingoverlapping echoes. That is, a second ping may be transmitted before allechoes from a first ping are received. This is possible as long as thetransmitted ping signals may be coded or otherwise distinguished suchthat echoes of a first ping may be recognized as distinct from echoes ofa second ping. Several coded excitation techniques are known to thoseskilled in the art, any of which may be used with a point-sourcemultiple aperture imaging probe. Alternatively, overlapping pings mayalso be distinguished by transmitting pings at different frequencies orusing any other suitable techniques. Using overlapping pings, evenhigher imaging frame rates may be achieved.

In some embodiments, prior to initiating an elastographic imagingprocess, an imaging window may be defined during a B-mode imagingprocess. The defined image window may be a section of the region ofinterest in which elastography is to be performed. For example, theimage window may be defined after any combination of probe positioning,depth-selection, zooming, panning, etc. In some embodiments, an imagewindow may be as large as an entire insonified region of interest. Inother embodiments, an image window may be only a smaller section of thecomplete region of interest (e.g., a “zoomed-in” section). In someembodiments, an image window may be defined after an imaging sessionusing echo data retrieved from a raw data memory device.

FIG. 7 illustrates an embodiment of an elastography process 80 using aprobe such as that shown in FIG. 1 . In the illustrated embodiment, anelastography process 80 may generally involve the steps of obtaining 82and storing 84 a baseline image, transmitting a shear-wave initiatingpulse (an init pulse) 86 into the region of interest 50, imaging theregion of interest 50 using a high frame rate imaging process 88, andsubtracting the baseline image 90 from each frame obtained during thehigh frame rate imaging process 88. The remaining series of “differenceframes” can then be analyzed to obtain information about the tissuedisplaced by the shear wave 56 propagating through the tissue of theregion of interest 50. The propagation speed of the shear wave 56 may beobtained through analysis of the perturbation of tissue in thetime-series of difference frames.

In some embodiments, while imaging a selected image window within aregion of interest with an elastography-enabled ultrasound probe, aninit line 22 (shown in FIG. 1 ) may be displayed on an ultrasound imagedisplay screen overlying an image of the target region. In someembodiments, the ultrasound imaging system may continuously image theregion of interest with a high resolution imaging process as discussedabove with reference to FIG. 5 . Alternatively, any other desiredultrasound imaging process may be used to obtain an image of the regionto be analyzed by an elastography process.

Once the probe 10 is in a desired orientation such that the init line 22intersects a desired target object or portion of the region of interest,an elastography depth may be selected, and an elastography process 80may be initiated. In some embodiments, an elastography depth may beselected by an operator via a suitable user interface action. In otherembodiments, an elastography depth may be selected automatically by anultrasound imaging control system. In some embodiments, an elastographyprocess may be initiated manually by an operator of the ultrasoundsystem. In other embodiments, an elastography process 80 may beinitiated automatically by an ultrasound system upon automaticidentification of a structure to be inspected.

As shown in the embodiment of FIG. 7 , an elastography process 80 usinga probe such as that shown in FIG. 1 (or any other suitably configuredprobe) may begin by obtaining 82 and storing 84 a baseline image of thetarget region of interest 50. In one embodiment, the baseline image maybe formed by obtaining a single frame using a high-frame-rate imagingprocess such as that described above. In such embodiments, a baselineimage may be formed by transmitting an imaging pulse from a singletransducer element Tx from a first of the lateral transducer arrays 14,16 (e.g., the right array 16), and receiving echoes on multiple elementsof the second of the lateral transducer arrays 14, 16 (e.g., the leftarray 14). In some embodiments, echoes from the transmit pulse may alsobe received by receive elements on the first transducer array (e.g. theright array 16). The baseline image may then be formed and stored 84 foruse in subsequent steps. In an alternative embodiment, the baselineimage may be obtained 82 using a high resolution imaging process such asthat described above.

After obtaining a baseline image 82, the init transducer array may beoperated to transmit a shear-wave initiating pulse 86 into the region ofinterest. An init pulse may be produced by any suitable devices andmethods as described above. In some embodiments, the shear waveinitiating pulse may be focused along a displayed init line 22, and maybe focused at a particular depth within the region of interest.

After an init pulse is transmitted 86, the system may begin imaging theregion of interest at a high frame rate 88 using the lateral imagingarrays 14, 16. In some embodiments, the high frame rate imaging processmay comprise the process described above with reference to FIG. 6 . Inone embodiment, the high frame rate imaging process may comprisetransmitting a series of transmit pulses from a single transmit apertureTx, and receiving echoes at a plurality of elements on at least onereceive aperture. In some embodiments, the high frame rate imaging 88may be performed by transmitting ultrasound pulses from the sametransmit element (or aperture) as that used in the step of obtaining abaseline image 82. In some embodiments, the high frame rate imaging maycontinue at least until propagation of the induced shear wave hasstopped or has progressed to a desired degree. A duration of highframe-rate imaging time may be calculated in advance based on anexpected minimum propagation speed and an image size. Alternatively, thehigh frame rate imaging 88 may be stopped upon detecting the shearwave's propagation at an extent of an imaging frame.

In some embodiments, forming a single frame during a high frame rateimaging process 88 may include combining image layers obtained fromechoes received at different receiving transducer elements. For example,separate images may be formed from echoes received by each individualtransducer element of a receive aperture to form a single improvedimage. Then, a first image produced by echoes received by all elementsof a first receive aperture may be combined with a second image producedby echoes received by all elements of a second receive aperture in orderto further improve the quality of the resulting image. In someembodiments, the image resulting from such combinations may then be usedas a single frame in the high frame rate imaging process 88. Furtherexamples of such image combining are described in U.S. patentapplication Ser. No. 13/029,907 referenced above.

In some embodiments, the baseline image may then be subtracted 90 fromeach individual frame obtained in the high frame rate imaging process88. For example, each pixel value of a single frame may be subtractedfrom the value of each corresponding pixel in the baseline image. Theimage resulting from such subtraction may be referred to as a“difference image” or a “difference frame.” The difference images thusobtained will include pixel values representing substantially only theshear waveform plus any noise.

In some embodiments, the steps of obtaining a baseline image 82,transmitting an init pulse 86 continuously imaging at a high frame rate88, and obtaining difference image frames 90 may be repeated as manytimes as desired. The difference images from such multiple cycles may beaveraged or otherwise combined in order to improve a signal to noiselevel.

The propagating shear waveform may be detected along lines transverse tothe direction of the init pulse (e.g., as shown in FIG. 1 ) by detectingperturbation (i.e., small changes in an otherwise ‘normal’ pattern) insubsequent difference frames. The speed of the shear wave's propagationmay be obtained by determining the position of the shear wave inmultiple image frames obtained at known time intervals.

In some cases, the perturbation caused by a propagating shear wave mayproduce a relatively disbursed image of the propagating wave front. Forexample, perturbation may appear in a difference frame as a specklepattern 92 such as that shown in FIG. 8 . An approximate center line 94of the point cloud 92 may be determined and treated as representative ofthe position of the propagating shear wavefront. In some embodiments, aline, curve or other path 94 may be fit to the point cloud 92 using anysuitable path fit algorithm. For example, in some embodiments anabsolute value of the difference frame may be calculated, and a localposition of the shear wave may be determined by averaging the positionof the nearest x points.

Ping Based Elastography

Ping Based Multiple Aperture Imaging (PMA) systems provide additionalmethods of creating tissue displacement. In some embodiments FIG. 9 , amultiple aperture probe 10 is configured to generate one or more shearwaves. For example, the multiple aperture probe can generate a firstshear wave 32 by selecting transmitters located in a first coherentsection 14 of a concave transducer array, then can generate a secondshear wave 31 by selecting transmitters located in a second coherentsection 16 of the concave transducer. In one example, the secondcoherent section may be physically separated from the first coherentsection, The propagating shear waves 32 and 31 generate point clouds 92along the wavefronts 94 and 93, respectively, which can be detectedand/or imaged by the multiple aperture imaging system. As shown in FIG.9 , multiple off axis shear waves transmitted to an organ 50 or tissueof interest 70 provides multiple data sets from around the entire tissueand even surrounding tissues. In other PMA embodiments, shear waveamplitudes can be significantly reduced as compared to linear phasedarray transmitters.

PMA transmitters used in any small grouping of elements in probe 10 onarrays 12, 14, 16 or any combination of elements therein provide enoughunfocused acoustic energy for imaging; therefore, when increasing theping to utilize a greater number transmit transducers having an outputof more acoustic energy, a much lower amplitude but significantlydifferentiated shear wave is created. This shear wave's effect ontissues of interest can be tracked using Young's modulus even though itwas created by simply using more elements in Ping Based MultipleAperture Imaging system.

Shear wave data transiting the organ 50 or tissue of interest 70 frommultiple off axis directions provides more data that can be averaged andweighted, but also provides better edge detection. Ping Base MultipleAperture probes can be constructed to provide up to 270 degrees of viewof the region of interest (extra corporeally) and 360 degree viewsintra-cavity, providing significant edge detection data from multipledifferent aspect angles.

Ultrasound shear waves typically result in only a few microns of tissuedisplacement. Since this amount is less than the resolution of mostimaging systems, detecting the displacement carries additionalchallenges. In some embodiments, tissue displacement induced by shearwaves may be detected in terms of the phase shift of the return ofB-mode imaging echoes.

In some cases, shear waves can cause perturbation on a tissue edge thatare defined by speckle patterns. These speckle patterns may appear inone or more frames of data as shown in FIG. 9 . The signal processor ofthe imaging system may detect/identify the speckle patterns along thetissue edge, which can be used for edge detection. In some embodiments,the tissue edge can be identified on a display of the ultrasound imagingsystem. For example, a line, marker, indicator, or other visual markingcan be overlaid onto the display to show the tissue edge(s). Anapproximate center line of the point cloud 92 may be determined andtreated as representative of the position of the propagating shearwavefront. Wavefronts along multiple axes may be measured simultaneouslyas represented by the initiation of second shear wave 31 after firstshear wave 32. The wavefronts 94 and 93 can traverse an area ofdiffering tissue density at tissue of interest 70. Note, each shear waveprovides a different force on the tissue of interest, and this force cantherefore be measured moving in multiple axes for more accurate densityinformation.

In some embodiments, a line, curve or other path may be fit to the pointcloud(s) 92 using any suitable path fit algorithm. For example, in someembodiments an absolute value of the difference frame may be calculated,and a local position of the shear wave may be determined by averagingthe position of the nearest points.

In some embodiments, the analysis may be limited to only a portion ofthe point cloud 92 (and/or a corresponding center line 94). For example,if it is determined (by visual inspection or by automated analysis) thata small segment of the shear wavefront is propagating faster thanadjacent segments, the region(s) of apparent higher or lower propagationspeed may be selected, and the speed of propagation may be calculatedfor only that portion of the shear wavefront. When adjacent shear waveshave identified edges through speckle noise patterns and highlightedwith fiducial marks, multiple propagation pathways can be analyzedsimultaneously. In some embodiments, an elastographic measurement isprovided for shear wave data resulting from the transmission from afirst coherent transmitter window solely along one axis. In someembodiments, shear wave data resulting from the transmission from afirst coherent transmitter window is measured, weighted or averagedalong multiple axes. In some embodiments, shear wave data is collectedand compared along one axis for two successive transmissions from afirst coherent transmitter window and a second coherent transmitterwindow that may be located and physically separated from the firstcoherent transmitter window. In some embodiments, shear wave data iscollected, weighted or averaged along multiple axis for two successivetransmissions from a first coherent transmitter window and a secondcoherent transmitter window that may be located and physically separatedfrom the first coherent transmitter window.

Referring to FIG. 8 , by calculating a distance between the focused initline 22 and the fit line 94 in a given difference frame, an approximateposition of the shear wave in the given difference frame may becalculated. Alternatively, referring to FIG. 9 , by calculating adistance between the focused init line 31 and the fit line 93 along aseparate axis in a given difference frame, an approximate position ofthe shear wave in the given difference frame may be calculated. The rateof propagation of the wavefront between any two frames may be determinedby dividing the distance traveled by the shear wave by the time thatelapsed between obtaining the two frames. In alternative embodiments,the position of a shear wave in any given frame may be measured relativeto any other suitable datum. In some embodiments, this process ofcalculating distance between frames can be used along multiple axessimultaneously.

In various embodiments, the number of frames needed to measure thepropagation speed of a shear wave may vary. In some embodiments anapproximate speed measurement may be obtained from as few as two orthree frames obtained at known time intervals. In other embodiments, atleast ten frames obtained at known time intervals may be needed toobtain a sufficiently accurate time measurement. In further embodiments,at least 100 frames obtained at known time intervals may be used toobtain a more accurate time measurement. In still further embodiments,200 frame or more may be used. Generally, the accuracy of shear wavepropagation speed measurements may increase with the number of framesfrom which such measurements are made. As the number of framesincreases, so does computational complexity, so the number of frames tobe used may be balanced with available processing capabilities.

When more than two frames are available to be used for measuringpropagation speed, any number of algorithms may be used. For example, insome embodiments the shear wave position may be detected in eachavailable frame, a speed may be calculated between each consecutive pairof frames, and the results of all such speed measurements may beaveraged to obtain a single speed value. In other embodiments, speedmeasurements may be calculated based on time intervals and relativeshear wave positions between different and/or variable numbers offrames. For example, propagation speed may be calculated between everythree frames, every five frames, every 10 frames, every 50 frames, etc.Such measurements may then be averaged with one another and/or withmeasurements obtained from consecutive frame pairs. Weighted averagesmay also be used in some embodiments.

In some embodiments, an entire elastography process 80 (FIG. 7 ) may berepeated at different focus depths relative to the init transducer array12. In some embodiments, un-beamformed elastography echo data obtainedat various depths may be stored and combined into a single 2D or 3D dataset for further post processing and/or for later viewing and analysis.In various embodiments, un-beamformed elastography echo data may becaptured and stored for later processing on the imaging system or anyother suitable computing hardware.

In alternative embodiments, the propagation speed of a shear wave may bemeasured by detecting the speed of moving/displaced tissues using themultiple aperture Doppler techniques described in Applicant's co-pendingU.S. patent application Ser. No. 13/690,989, filed Nov. 30, 2012, titled“Motion Detection Using Ping-Based And Multiple Aperture DopplerUltrasound.”

Once the shear wave is captured and its propagation speed is measured,the hardness of the tissue in the region of interest, as quantified byYoung's modulus (E) can be measured or determined by a controller,signal processor or computer. Elasticity (E) and shear wave propagationspeed (c) are directly related through the simple formula:

E=3ρc ²

Where ρ is the density of tissue expressed in kg/m³. Because the densityof tissues tends to vary minimally, an approximate density value may beassumed for the purpose of calculating elasticity using a measuredpropagation speed value. The fact that the speed term is squared furtherminimizes the effect of any error in the assumed density value. Thus,the elasticity of the tissue may be calculated after measuring only theshear wave propagation velocity c and using an assumed approximate valuefor tissue density.

In some embodiments, the density value may be stored in a digital memorydevice within or electronically accessible by the controller. In otherembodiments, the density value may be manually entered or edited by auser via any suitable user interface device. Once the speed of shearwave propagation has been measured for a desired area within the regionof interest, the controller may retrieve the density value and calculatethe elasticity for the desired area.

In some embodiments, elasticity estimates may be overlaid on an image ofthe region of interest. In some embodiments, such an overlay may beprovided as a color coded shaded image, showing areas of high elasticityin contrasting colors to areas of relatively low elasticity.Alternatively, a propagating shear wave may be displayed on an image. Insome embodiments, a propagating shear wave may be displayed as ananimated moving line, as changing colors, as a moving point cloud or inother ways. In further embodiments, a numeric value of a shear wavepropagation speed may be displayed. In other embodiments, numeric valuesof elasticity may be displayed on an image of the region of interest.Soft tissues will tend to have relatively small values of elasticity,and liquid-filled areas do not conduct shear waves at all.

Some embodiments will designate a controllable region of interest on thedisplay. Elastography values that contain common values within theregion of interest will be grouped together and can be colorized. Insome embodiments, specific types of tissue (e.g. bone, blood, lung) canbe assigned specific colors based on known elastographic values. Wheregroupings of tissue differ by a statistical value, for example greaterthan or less than 10%, an artificial demarcation will be drawn betweenvisually presented groups in the region of interest. This statisticalvalue is selectable, and will vary in its effective presentation oftissue types based on the tissue being evaluated. For instance, softplaques located in an artery may require a very low differentiationvalue, whereas a kidney stone in the bile duct may require a highdifferentiation value. Once implemented, the line indicating demarcationis presented in white, but specific colors for the demarcation may beselected.

Ping Based Multiple Aperture Imaging Edge Detection

Edge detection is normally an image processing or post processingtechnique used in medical imaging where a mathematical method is used indiscerning where brightness changes sharply within a beamformed orcomputed image area. In Ping Based Multiple Aperture (PMA) Ultrasound,edges can also be determined by tracking visual and sub-visualinterference patterns sometimes called speckle noise patterns. Whenstruck by a wavefront, reflectors and scatterers will tend to re-radiatea wavefront in a direction generally dictated by physics.

In FIG. 10 , an unfocused wavefront 71 emanating from transducerelements of a first coherent section 14 of the probe 10 pass throughunknown cyst 70 inside a larger organ 65. Some of the wavefront 71continues on through the unknown cyst and into the organ. Some of thetransmission 71 is also reflected off of the unknown cyst 70 in awavefront 73. Wavefronts reacting to a change in tissue density, forexample the edge of an organ (e.g. the liver capsule) the similar effectof re-radiating wavefronts in patterns consistent with the effect of notjust a single scatterer, rather a line of scatterers associated with thetissue's edge. The amplitude of this reflected wavefront is translatedby receiver transducer into data that can be presented in an image. Thisvisual data is what is commonly used to create an edge.

Visual presentation of the tissue's edge is common today. Oftenpost-processing techniques, such as image compounding and persistence,are used to provide further visual clarity on where the tissue edge maybe.

The transmission from a ping-based imaging system is an unfocused wave.This transmission wavefront is affected by the tissues it transits onthe way to the region of interest. Like any pressure wave, the wavefrontis affected by the density of the tissues it is attempting to passthrough. Some tissues are more dense and the wave either accelerates orbounces off, some tissues are less dense either bounces off or slowsdown. The effects of these wavefront changes also aren't necessarilyaligned with the transmitter. For example, some of the wavefront willtravel in a new direction altogether. So when wavefront 71 crosses point66, the actual wave(s) crossing point 66 (if examined undermagnification) would look like 550 a from FIG. 11A. This representationlooks like a “fingerprint” associated with that point. For the period ofseveral milli-seconds, it follows that the tissues surrounding point 66will remain consistent, and not be affected by standard physiologicalfunctions such as breathing, blood flow or peristalsis for instance.Therefore, the fingerprint for point 66 will remain identifiable overdozens transmit cycles and can be identified by the ping-based imagingsystem. Similarly, the actual wave crossing point 67 could look like 550b from FIG. 11B. Additional transmission wavefronts crossing point 66from multiple transmitters used in a ping-based multiple apertureimaging system would add additional wavefront crossings to the data setat that point, and subsequently further identifying features.

Conventional phased array and plane wave systems are unable to achievethis unique identifying feature because the wavefront strength is bothsteered and too great. Wavelets work together to create a wavefront andtherefore aren't as pliable when addressing different tissue types.

Ping-based multiple aperture imaging systems, can therefore provide amore discernable edge to tissue identification, and an edge that can betracked over time. That is, once a tissue edge is identify by multiplecomputer assigned fiducial marks (fingerprints), then those marks can betracked over time. The process of using PMA systems to identify fiducialmarks and fingerprints is further described in U.S. Pat. No. 10,380,399.In essence, a fingerprint that is being tracked for several cycles then“hands off” its location to a new fingerprint that should remain inplace for several dozen more cycles, and so on. Ping-based multipleaperture imaging systems track these markings continuously. This can bevery useful when a therapy is being applied. For instance, when tissueis being biopsied, it may bulge or squirt out of the way of the needle;and, while a surgeon may be able to react in the moment to visual cues,a surgical robot performing the same procedure may not. The robot needstracking data, not visual data, in order to stay centered on the biopsytarget. Ping-based imaging systems can provide both visual cues andsub-visual fiducial tracking. Similar cases could be argued for oblationtherapies, surgical removal of tissue, measurement of arterial wallsize, measurement of biologic build up on artificial valves and manyother applications where visual presentation of edge data is notdetectable by a human operation without full magnification.

The effect of linear and curvilinear wavefronts from one set ofreflectors crossing other wavefronts from other reflectors can become apredictable matter of fact in a specific area of interest over smallperiods of time (e.g. less than one second). Transmitters using a 3.5MHz ping-based pulse would typically generate wavefronts that have amaximum wavelength of 0.44 mm. The crossing location of wavefronts wouldtherefore be only a small portion of the maximum amplitude, perhaps 0.1mm, and therefore not made visible on in the resolution of aconventional ultrasound system monitor. However, the act of thewavefronts crossing is occurring in a regular fashion and is detectableby a Ping Based Multiple Aperture Imaging system. Such a wavefrontcrossing is made consistent by the consistent types of tissue in theimmediate area of interest (often less than 1 mm square) and thereforemaking the crossing pattern of the wavefronts consistent over asignificant period of transmit cycles. Wavefronts and associatedcrossing wavefronts emanating from a tissue barrier at specificlocations in the region of interest can further be used to mark, outlineor track patterns associated with tissue normally outside the visualresolution of the imaging monitor. Examples such as individualcapillaries, microscopic calcium deposits or even individual blood cellsare typically below the imaging resolution provided by conventionalultrasound systems, but identifiable and available to Ping BasedUltrasound Imaging systems. Such structures are hereby being referred toas sub-visual.

A fiducial region may be defined manually by a human user orautomatically by an imaging system or other computing system. The stepof defining a fingerprint within the fiducial region may comprisedefining a sub-region within the fiducial region (manually orautomatically), and then defining the fingerprint as a collection ofdata samples from among a complete data set representing the entiresub-region. Alternately, the step of defining a fingerprint within thefiducial region may comprise directly defining a collection of datasamples from among a complete data set representing the entire fiducialregion.

In some embodiments, a fingerprint may be sized to be smaller than thelateral and/or axial resolution limits of the imaging system. In suchembodiments, the fingerprint point may represent a pattern oftransmitted wavefronts at a particular point within or on the objectrather than actual visible features of the object. In such embodiments,tracking of the fingerprint point may be improved by an imaging systemthat produces wavefronts that intersect at significant angles to oneanother at the object to be tracked. For example, if wavefrontsintersect at approximately 90 degrees to one another at the fingerprintpoint, a beamformed representation of the pattern of intersectingwaveforms may be substantially unique within at least the local regionof the insonified field compared with other similarly situated regions.In some embodiments, such a representation may comprise a cross-hatchedpattern of intersecting lines when displayed (whether or not such pointsare actually displayed).

Edge detection is imperative in understanding imaging. Edges enablemeasurement, enable motion tracking, and tissue differentiation. Incardiology, edge detection and tracking enable measurement of stenosis,aneurism, valve movement and function, and heart wall movement, all arecritical functions. In oncology, edge detection enable cystic lesionidentification, size, density, measurement before and after therapeutictreatment are all essential. In imaging guidance, accurate outlines oftissue edges during treatment enables oblation to be conducted in realtime. More difficult guidance procedures, such as determining thelocation of a lesion inside bone for biopsy uses fiducial markers insidethe bone that can then be tracked on a PMA system.

The ability to identify, measure and track sub-visual tissue edgesprovides substantial advantages to both diagnosticians and therapists intoday's computer aided environment. In liver tissue oblation procedures,physicians must outline the region to be ablated. A margin of safety isadded beyond what can be marked visually. This margin of safety is addedprecisely because visual indications can often miss sub-visual tissuesthat are intended to be ablated. Sub-visual tissue demarcation thereforebecomes more valuable in reducing the amount of good tissue that isablated in a procedure.

Multiple Aperture Elastography, or utilizing shear wave information frommultiple axes can provide tissue characterization throughout the regionof interest. Conventional elastography provides a simple density oftissue within a larger organ. Multiple aperture elastography provides 3Delastography even outside of the primary organ and into surroundingtissues. Enabling examples of tissue density in multiple organs andglands at the simultaneously. In oncology, for instance, a ping basedmultiple aperture imaging system may examine the liver, kidney andpancreas at the same time.

In some embodiments, the phase shift of an unfocused wave can bemeasured by tracking reflected wavefronts in fingerprint patternsassociated with edges located in the tissue. In FIG. 10 , an unfocusedwavefront 71 transmitted by a first coherent section 14 of a concaveprobe 10 reflects in all directions off of scatterers from tissuecontaining a different density and associated speed-of-sound in theregion of interest 50. Therefore, each speed-of-sound differentiatedbarrier creates a waveform or fingerprint pattern associated with eachtissue edge. Transducer elements within a first coherent section 14 caninitiate an unfocused transmit wavefront 71, which creates reflectedfingerprint wavefront patterns 60 (e.g., shown as polka dot patterns) asit progresses through each tissue type, such as for example, bone 61,anterior organ capsule 62, anterior, medial, lateral and posteriorcystic structure 63, and posterior organ capsule 62.

A ping transmission from a second coherent section 16 of the concaveprobe 10 can cause wavefront 72 and subsequent fingerprint patterns 65(e.g., illustrated as x-shaped patterns) associated with those samescatterers 61, 62, and 63. However, the offset angle between the firstcoherent section 14 and the second coherent section 16 can providedifferentiated data for wavefront fingerprint pattern recognition.Speckle patterns located near a single reflector can be outlined andassigned fiducial marks or edges. PMA ultrasound, however, does notrequire the visual beamforming of an image in order to create fiducialsmarks in edge location.

In some embodiments a fingerprint may be sized to be smaller than thelateral and/or axial resolution limits of the imaging system. In suchembodiments, the fingerprint point may represent a pattern oftransmitted wavefronts at a particular point within or on the objectrather than actual visible features of the object. In such embodiments,tracking of the fingerprint point may be improved by an imaging systemthat produces wavefronts that intersect at significant angles to oneanother at the object to be tracked. For example, if wavefrontsintersect at approximately 90 degrees to one another at the fingerprintpoint, a beam-formed representation of the pattern of intersectingwave-forms may be substantially unique within at least the local regionof the insonified field compared with other similarly situated regions.In some embodiments, such a representa-tion may comprise a cross-hatchedpattern of intersecting lines when displayed (whether or not such pointsare actu-ally displayed).

FIG. 11A and FIG. 11B illustrate some example cross-hatched fingerprintpatterns 550 a, 550 b derived from data collected using a multipleaperture ping-based ultrasound imaging system as described herein. Thepatterns 550 a, 550 b are shown in exaggerated black-and-white in orderto highlight the pattern. In an imaging system, patterns may be producedin grayscale and may therefore provide more nuanced detail.

Fingerprint patterns can automatically located near tissue edges andtherefore tracked. For instance, referring to FIG. 10 , fingerprintpatterns can be identified at locations 66 and immediately adjacentfiducial mark 67. It follows then that fingerprint identification ofidentical fingerprint patterns at all locations of along the edges ofthe tissue could therefore be identified and tracked.

In some embodiments, edges are created and fiducial marks assigned basedon the patterns identified in data strings in memory. In otherembodiments, images are beamformed and sub-visual patterns areidentified around reflectors, edges defined and fiducial marks assigned.For example, referring to FIG. 10 , the polka dot fingerprint patterns60 associated with the unfocused transmit 71 can be identified andtracked, and similarly, the x-shaped fingerprint patters 65 associatedwith the unfocused transmit 72 can be identified and tracked. Thefingerprint patterns associated with different unfocused transmissionscan be used to identify multiple edges of a target tissue. For example,referring to FIG. 10 , the fingerprint patterns associated with theunfocused transmit 71 can be used to identify the top and bottom edgesof target tissue 70, while the fingerprint patterns associated withunfocused transmit 72 can be used to identify the left and right edgesof target tissue 70. Together, the various identified fingerprintpatterns can be used to identify some or all edges of a target tissue,such as a tumor or cyst.

Raw Echo Data Memory

Various embodiments of the systems and methods described above may befurther enhanced by using an ultrasound imaging system configured tostore digitized echo waveforms during an imaging session. Such digitalecho data may be subsequently processed on an imaging system or on anindependent computer or other workstation configured to beamform andprocess the echo data to form images. In some embodiments, such aworkstation device may comprise any digital processing system withsoftware for dynamically beamforming and processing echo data using anyof the techniques described above. For example, such processing may beperformed using data processing hardware that is entirely independent ofan ultrasound imaging system used to transmit and receive ultrasoundsignals. Such alternative processing hardware may comprise a desktopcomputer, a tablet computer, a laptop computer, a smartphone, a serveror any other general purpose data processing hardware.

In various embodiments, received echo data (including echoes receivedduring a high frame rate imaging process) may be stored at variousstages from pure analog echo signals to fully processed digital imagesor even digital video. For example, a purely raw analog signal may bestored using an analog recording medium such as analog magnetic tape. Ata slightly higher level of processing, digital data may be storedimmediately after passing the analog signal through an analog-to-digitalconverter. Further processing, such as band-pass filtering,interpolation, down-sampling, up-sampling, other filtering, etc. may beperformed on the digitized echo data, and raw data may be stored aftersuch additional filtering or processing steps. Such raw data may then bebeamformed to determine a pixel location for each received echo, therebyforming an image. Individual images may be combined as frames to formvideo. In some embodiments, it may be desirable to store digitized echodata after performing very little processing (e.g., after some filteringand conditioning of digital echo data, but before performing anybeamforming or image processing). Some ultrasound systems storebeamformed echo data or fully processed image data. Nonetheless, as usedherein, the phrases “raw echo data” and “raw data” may refer to storedecho information describing received ultrasound echoes (RX data) at anylevel of processing prior to beamforming. Raw echo data may include echodata resulting from B-mode pings, Doppler pings, or any other ultrasoundtransmit signal.

In addition to received echo data, it may also be desirable to storeinformation about one or more ultrasound transmit signals that generateda particular set of echo data. For example, when imaging with a multipleaperture ping ultrasound method as described above, it is desirable toknow information about a transmitted ping that produced a particular setof echoes. Such information may include the identity and/or position ofone or more a transmit elements as well as a frequency, magnitude, pulselength, duration or other information describing a transmittedultrasound signal. Transmit data is collectively referred herein to as“TX data”.

In some embodiments, TX data may also include information defining aline along which a shear-wave initiating pulse is transmitted, andtiming information indicating a time at which such a shear-waveinitiating pulse is transmitted relative to received echo data. In someembodiments, such TX data may be stored explicitly in the same raw datamemory device in which raw echo data is stored. For example, TX datadescribing a transmitted signal may be stored as a header before or as afooter after a set of raw echo data generated by the transmitted signal.

In other embodiments, TX data may be stored explicitly in a separatememory device that is also accessible to a system performing abeamforming process. In embodiments in which transmit data is storedexplicitly, the phrases “raw echo data” or “raw data” may also includesuch explicitly stored TX data. In still further embodiments, transducerelement position information may be explicitly stored in the same or aseparate memory device. Such element position data may be referred to as“calibration data” or “element position data”, and in some embodimentsmay be generally included within “raw data.”

TX data may also be stored implicitly. For example, if an imaging systemis configured to transmit consistently defined ultrasound signals (e.g.,consistent magnitude, shape, frequency, duration, etc.) in a consistentor known sequence, then such information may be assumed during abeamforming process. In such cases, the only information that needs tobe associated with the echo data is the position (or identity) of thetransmit transducer(s). In some embodiments, such information may beimplicitly obtained based on the organization of raw echo data in a rawdata memory. For example, a system may be configured to store a fixednumber of echo records following each ping. In such embodiments, echoesfrom a first ping may be stored at memory positions 0 through ‘n’ (where‘n’ is the number of records stored for each ping), and echoes from asecond ping may be stored at memory positions n+1 through 2n+1. In otherembodiments, one or more empty records may be left in between echo sets.In some embodiments received echo data may be stored using variousmemory interleaving techniques to imply a relationship between atransmitted ping and a received echo data point (or a group of echoes).Similarly, assuming data is sampled at a consistent, known samplingrate, the time at which each echo data point was received may beinferred from the position of that data point in memory. In someembodiments, the same techniques may also be used to implicitly storedata from multiple receive channels in a single raw data memory device.

In some embodiments, raw TX data and raw echo data may be captured andstored during an imaging session in which an elastography process isperformed. Such data may then be later retrieved from the memory device,and beamforming, image processing, and shear-wave speed measurementsteps may be repeated using different assumptions, inputs or algorithmsin order to further improve results. For example, during suchre-processing of stored data, assumed values of tissue density orspeed-of-sound may be used. Beamforming, image layer combining, or speedmeasurement averaging algorithms may also be modified during suchre-processing relative to a real-time imaging session. In someembodiments, while reprocessing stored data, assumed constants andalgorithms may be modified iteratively in order to identify an optimumset of parameters for a particular set of echo data.

Automated Tissue Characterization, Edge Detection and Tracking UsingPing-Based Multiple Aperture Imaging

A benefit of using off axis shear waves inside the same data setcollected by a ping-based multiple aperture imaging system is theability to identify tissue edges with higher certainty. Tissue edgesemanating from the naturally occurring groupings of common tissuedensities, enables the system to visually present lines of demarcationbetween tissue types. Once lines of demarcation between tissue types arebe presented, a ping-based imaging system can then track the edges oftissues presented in the region of interest without manual placement bythe operator.

FIG. 12 is a flowchart 600 that describes the operation of automatedtissue characterization, edge detection and tracking using ping-basedmultiple aperture imaging and shear waves. The region of interest isimaged using a multiple aperture probe of a ping-based ultrasound systemultrasound in step 602. In step 604, a shear wave can be introduced fromone of a plurality of coherent transmit windows along one axis from themultiple aperture probe. At step 606, data is collected and tissuedensities calculated and stored in memory. Data can be utilized at thispoint and move onto selecting tissue differentiation values. However, insome embodiments, as shown by step 608, shear waves are firedsequentially from multiple different axes along the probe, datacollected sequentially for each of those axes, and aggregated.

After a sequence of shear wave data collections, at step 610 theoperator or the ping-based ultrasound system can select a tissuedifferentiation value. This value can act as a statistical filterbetween tissue density types. At step 612, such tissue densities insideof the differentiation value can be grouped by the ping-based ultrasoundsystem. Tissues of common values can be assigned a common color fortissue characterization purposes, and the various colors representingdifferent tissues can be displayed or inserted into ultrasound images bythe ping-based ultrasound system. Where tissue is differentiated, ademarcation line can be inserted into the ultrasound images (and/ordisplayed) at step 614. Often this demarcation line is white or black,but can be automatically selected to standout from the most commondisplayed colors.

With known lines of demarcation being presented in a single pixel orvoxel absolute color, lines of demarcation can be inferred. At step 616,fiducial marks can be assigned by the computer to be on either side oflines of demarcation. These fiducial marks can then be tracked in step618. Even if the visual line of demarcation is lost, the actual line oftissue density differentiation can still be marked by the fiducialmarker, and lines can be updated in step 620. It is common forbreathing, pulsing arteries, peristalsis or the operator moving theprobe to change the tissue location inside the viewing region ofinterest. This can be dangerous when a therapy, or intervention istaking place such as a needle biopsy or intravenous device implantation.Automated tracking of tissue edges may reduce risk from theseprocedures.

FIG. 13 is a flowchart 700 that describes the operation of automatedtissue characterization, edge detection and tracking using ping-basedmultiple aperture imaging. Referring to step 702 of flowchart 700, themethod can include initiating imaging of a region of interest with oneor more unfocused ping transmissions from a first coherent transmitterof a multiple aperture probe of a ping-based ultrasound system. In someimplementations, the initiating imaging step can include transmittingone or more unfocused ping transmissions with a greater number oftransducers and/or acoustic energy than is normally used for ping-basedultrasound imaging. For example, if ping-based imaging typicallyrequires 1-10 transducers of a first coherent transmitter, the methodcan include transmitting unfocused ping transmissions from moretransducers, such as 10-20 transducers or 10-50 transducers of a firstcoherent transmitter. Alternatively, the step can include transmittingpings at a higher amplitude (i.e., to produce more acoustic energy) thanis typical when imaging with a ping-based system.

At steps 702 and 704 of flowchart 700, the method can include receivingecho data from the unfocused ping with first and second receivetransducers. At step 708, the method can further include combining thereceived echo data from the first and second receive transducers.

At step 710, the method can include identifying speckle noise patternsassociated with tissue edges, or associated with changes in tissuedensity or elasticity. As described in step 702, the method can includetransmitting pings from more transducers than are normally used forping-based imaging, or alternatively, transmitting more acoustic energyinto the tissue. These increased ping waveforms (or increased amplitude)can result in speckle noise patterns developing at or around tissueedges within the region of interest. In some implementations, normalpings used for ultrasound imaging can also result in speckle noisepatterns that can be identified by the ultrasound system. The ping-basedultrasound system can then detect and/or identify these speckle noisepatterns.

At step 712, fiducial markers can be manually or automatically (i.e., bythe ping-based ultrasound imaging system) assigned to the tissue edgesor changes in tissue density/elasticity. In one implementation, thefiducial markers are assigned directly to the tissue edges. In anotherimplementation, the fiducial markers can be assigned to inner/outeredges of the tissue edges.

At step 714, the ultrasound system can form an image of the region ofinterest. As described above, the image can be formed with theping-based ultrasound system and can include transmitting an unfocusedping or pings from a transmitter, receiving echoes on at least first andsecond receive arrays, combining the received echoes, and forming theimage from this combined data.

At step 716, the fiducial marks from step 712 can be assigned, overlaid,or applied to the image from step 714. This image with the fiducialmarks can be stored as a baseline image.

Next, at step 718, the region can be continuously imaged with theping-based ultrasound imaging system. For example, a plurality ofunfocused pings can be transmitted by the transmitter into the region ofinterest, and echo data can be received, stored, and combined by theultrasound system. In one implementation, the pings that are transmittedduring imaging have lower amplitude or acoustic energy than the pingsthat were used at step 702. In other implementations, the pings are thesame or similar to the pings from step 702.

At step 720, movement of the fiducial marks can be measured for each newframe of the imaging. Thus, as the region of interest is continuouslyimaged in step 718, the pings moving through the region of interest cancause the tissue edges, and therefore the fiducial marks associated withthe tissue edges, to move. This movement can be detected and measured bythe ultrasound system.

Finally, at step 722, the ultrasound system can be configured to computethe tissue density of tissues in the region of interest. For example,the ultrasound system can use the movement of the tissue edges in step720 to compute the tissue density. As previously discussed, the rate ofpropagation of the wavefront for a fiducial mark between any two framesmay be determined by dividing the distance traveled the mark by the timethat elapsed between obtaining the two frames. In alternativeembodiments, the position of a fiducial mark moved in any given framemay be measured relative to any other suitable datum. In someembodiments, this process of calculating distance between frames can beused along multiple axes simultaneously.

Once the fiducial edge movement and its propagation speed is measured,the hardness of the tissue in the region of interest, as quantified byYoung's modulus (E) can be measured or determined by a controller,signal processor or computer.

For ease of presentation, all tissue not of the same density can besubtracted out while stronger ping based waves are introduced into themedium, and tissue edges with fiducial markers are easily tracked.

Although this invention has been disclosed in the context of certainpreferred embodiments and examples, it will be understood by thoseskilled in the art that the present invention extends beyond thespecifically disclosed embodiments to other alternative embodimentsand/or uses of the invention and obvious modifications and equivalentsthereof. Various modifications to the above embodiments will be readilyapparent to those skilled in the art, and the generic principles definedherein may be applied to other embodiments without departing from thespirit or scope of the invention. Thus, it is intended that the scope ofthe present invention herein disclosed should not be limited by theparticular disclosed embodiments described above, but should bedetermined only by a fair reading of the claims that follow.

In particular, materials and manufacturing techniques may be employed aswithin the level of those with skill in the relevant art. Furthermore,reference to a singular item, includes the possibility that there areplural of the same items present. More specifically, as used herein andin the appended claims, the singular forms “a,” “and,” “said,” and “the”include plural referents unless the context clearly dictates otherwise.As used herein, unless explicitly stated otherwise, the term “or” isinclusive of all presented alternatives, and means essentially the sameas the commonly used phrase “and/or.” Thus, for example the phrase “A orB may be blue” may mean any of the following: A alone is blue, B aloneis blue, both A and B are blue, and A, B and C are blue. It is furthernoted that the claims may be drafted to exclude any optional element. Assuch, this statement is intended to serve as antecedent basis for use ofsuch exclusive terminology as “solely,” “only” and the like inconnection with the recitation of claim elements, or use of a “negative”limitation. Unless defined otherwise herein, all technical andscientific terms used herein have the same meaning as commonlyunderstood by one of ordinary skill in the art to which this inventionbelongs.

A complete listing of the claims follows:
 1. A method of identifyingtissue edges with ultrasound imaging, comprising the steps of:transmitting a first unfocused ultrasound pulse into a tissue region ofinterest including one or more tissue edges; transmitting a secondunfocused ultrasound pulse into the tissue region of interest; receivingechoes of the second unfocused ultrasound pulse; identifying one or morespeckle noise patterns associated with the one or more tissue edges inthe received echoes; assigning fiducial markers to the one or moretissue edges; transmitting a third unfocused ultrasound pulse into thetissue region of interest; measuring a movement of the fiducial markers;and computing a tissue density of at least one tissue within the tissueregion of interest.
 2. The method of claim 1, further comprising formingan image of the tissue region of interest with the received echoes andthe fiducial markers.
 3. The method of claim 1, wherein the firstunfocused ultrasound pulse has greater acoustic energy than the secondunfocused ultrasound pulse.
 4. The method of claim 3, wherein the firstunfocused ultrasound pulse is transmitted with more ultrasoundtransducers than the second unfocused ultrasound pulse.
 5. The method ofclaim 1, wherein the speckle noise patterns are caused by reflections ofthe first unfocused ultrasound pulse from the one or more tissue edges.6. The method of claim 1 wherein the movement comprises a distancetraveled by the fiducial markers in response to the third unfocusedultrasound pulse.
 7. The method of claim 1 wherein the movementcomprises a propagation speed of the fiducial markers in response to thethird unfocused ultrasound pulse.
 8. An ultrasound imaging system,comprising: a first ultrasound transducer array configured to transmit awavefront that induces a propagating shear wave in a region of interest;a second ultrasound transducer array configured to transmit circularwaveforms into the region of interest and receive echoes of the circularwaveforms; and a signal processor configured to form a plurality ofB-mode images of the region of interest from the circular waveforms, thesignal processor being further configured to identify one or morespeckle patterns along a tissue edge caused by the propagating shearwave to identify the tissue edge.
 9. The system of claim 8, wherein thefirst ultrasound transducer array comprises an array of phased-arrayelements.
 10. The system of claim 8, wherein the first ultrasoundtransducer array comprises an annular array of piezoelectric rings, andthe signal processor is further configured to focus the wavefront atvarious depths by adjusting phasing delays.
 11. The system of claim 10,wherein the first ultrasound transducer array comprises a switched ringtransducer.
 12. The system of claim 1 wherein the first ultrasoundtransducer array comprises a single piezoelectric transducer.
 13. Thesystem of claim 8, wherein the frame rate is at least 500 fps.
 14. Thesystem of claim 8, wherein the frame rate is at least 1,000 fps.
 15. Thesystem of claim 8, wherein the frame rate is at least 2,000 fps.
 16. Thesystem of claim 8, wherein the frame rate is at least 4,000 fps.
 17. Thesystem of claim 8, wherein the signal processor is configured toidentify the propagating shear wave as a point cloud moving through theregion of interest.
 18. The system of claim 8, wherein the signalprocessor is configured to define an image window identifying a sectionof the region of interest with a combination of zooming, panning, anddepth selection.
 19. The system of claim 18, wherein the system isconfigured to display a contemporaneous B-mode image of a selected imagewindow.
 20. A method of identifying tissue edges with ultrasound, themethod comprising the steps of: forming a baseline image of a region ofinterest with an ultrasound imaging system; transmitting an ultrasonicpulse configured to induce a propagating shear wave in the region ofinterest; transmitting a plurality of unfocused ultrasound pings intothe region of interest; forming a plurality of image frames of theregion of interest from received echoes of the plurality of unfocusedultrasound pings; subtracting the baseline image from at least two ofthe formed image frames to obtain at least two difference frames;identifying one or more speckle patterns along a tissue edge in theregion of interest caused by the propagating shear wave. 21.-27.(canceled)